Nuclear medical diagnostic device

ABSTRACT

Parameters T 1 , T 2 , and K required by a scintillator array identification mechanism in a two-stage scintillator γ-ray detector (depth of interaction (DOI)) are accurately and easily determined. The parameters required by the scintillator array identification mechanism are determined with reference to a first signal count ratio, which is obtained by irradiating a γ-ray on each scintillator array with luminescence pulses in an incident depth direction of the γ-ray having different attenuation time during an inspection stage of the γ-ray detector single unit. Furthermore, a second signal count ratio is obtained by irradiating the γ-ray on a front surface of the γ-ray detector single unit, and then a third signal count ratio is obtained by irradiating the γ-ray on the front surface after the γ-ray detector single unit is installed in a PET device.

BACKGROUND OF THE INVENTION

1. Field of the Invention

The present invention generally relates to a nuclear medical diagnosticdevice (an emission computed tomography (ECT) device), in which anradioactive agent is applied to a test subject, and a pair of γ-raysdischarged by positron radioactive isotopes (RIs) accumulated in atarget portion of the test subject is measured concurrently, so as toobtain a tomogram of the target portion. In particular, the presentinvention relates to a technique of counting γ-rays at the same time.

2. Description of Related Art

A positron emission tomography (PET) device is taken as an example toillustrate a nuclear medical diagnostic device, that is, an ECT device.The PET device is formed in the following manner. Opposite γ-raydetectors are used to detect two γ-rays at an angle of approximately180° discharged from a target portion of a test subject. When the γ-raysare detected (counted) at the same time, a tomogram of the detected bodyis formed again. Further, some of the γ-ray detectors used to count theγ-rays at the same time in the PET device are formed by scintillatorsand photomultipliers. The scintillators then emit light after the γ-raysdischarged from the test subject are incident thereon, and thephotomultipliers convert the light emitted by the scintillators to anelectric signal.

In principle and in most cases, the γ-rays discharged from a visualfield centre are obliquely incident on the scintillators of the γ-raydetector D as shown in FIG. 16. When the divided scintillator is notpresent in the γ-ray incident direction, the γ-ray is detected in boththe correct positions and incorrect positions. That is, the visualdifference error becomes larger from the visual field centre to theperipheral parts, such that the tomogram obtained by the PET devicebecomes inaccurate.

Therefore, as shown in FIG. 17, a γ-ray detector is provided. In theγ-ray detector, the scintillators are divided (optically combined) intoscintillators having the luminescence pulses in the γ-ray incidentdirection with different attenuation times. For example, when a γ-raydetector MD is used, in which the scintillators are divided into ascintillator array with a shorter γ-ray attenuation time on a γ-rayincident side and a scintillators array with a longer γ-ray attenuationtime on a photomultiplier side, the positions of the discharged γ-rayscan still be detected accurately even if the γ-rays are obliquelyincident on the scintillators of the γ-ray detector MD. A more accuratetomogram is resulted, and improvement is achieved (for example, pleaserefer to Patent Reference 1 and 2).

Further, the specific mechanism for detecting the γ-ray position withrespect to the scintillator array with the shorter attenuation time andthe scintillator array with the longer attenuation time, stacked in theγ-ray incident direction, includes the following mechanisms: an addingmechanism for converting an electric signal output from a lightreceiving element, that is, converting an analog signal S_(F) (thesignal of the scintillator array with the shorter attenuation time) orS_(R) (the signal of the scintillator array with the longer attenuationtime) to a digital signal by using an A/D converter as shown in FIG. 18,and adding sequentially the digital signals converted by the A/Dconverter as shown in FIG. 19; an identification value calculatingmechanism for calculating an identification value representing a valueA_(T1)/A_(T2) or B_(T1)/B_(T2) obtained by dividing an intermediateadded value A_(T1) or B_(T1) by a total added value A_(T2) or B_(T2), inwhich the intermediate added value A_(T1) or B_(T1) is a value obtainedby adding the digital signals during a period starting from the point atwhich the luminescence of the luminescence pulse is emitted from thescintillator block to an intermediate point of the period at which theluminescence ends, that is, to an intermediate moment in the addingmechanism, and the total added value A_(T2) or B_(T2) is a valueobtained by adding the digital signals during a period starting from thepoint at which luminescence of the luminescence pulse is emitted fromthe scintillator block to point at which the luminescence ends, in theadding mechanism; a mechanism for deciding a medium value K according toa maximum value and a minimum value in the identifying values calculatedby the identification value calculating mechanism; and a determinationmechanism for determining whether an identification value calculated bythe identification value calculating mechanism is greater or smallerthan the medium value K.

Moreover, the existing nuclear medical diagnostic device determines theparameters used for judging in the following manner. That is, in thecase of a two-stage scintillator detector 112, for example, havingscintillator arrays of a two-stage structure as shown in FIG. 20,parameters T₁, T₂, and K required by the scintillator arrayidentification mechanism are decided in the following manner. In aninspection stage of the γ-ray detector single unit, the two-stagescintillator detector 112 disposed in a dark box 115 inputs theparameter of an initial value to a processing circuit for inspection,starts to irradiate γ-ray on a front surface 110 of the scintillatorarray, and then calculates a signal count N₁ and a signal count N₂through a determination calculation, in which the signal count N₁ isdetermined as the count of signals from the front surface 110 of thescintillator array, and the signal count N₂ is determined as the countof signals from a rear surface 111 of the scintillator array.Thereafter, as shown in FIG. 21, the γ-ray is only irradiated on therear surface 111 of the scintillator array, and a signal count N₂′ and asignal count N₁′ are calculated through a determination calculation, inwhich the signal count N₂′ is determined as the count of signals fromthe rear surface 111 of the scintillator array, and the signal count N₁′is determined as the count of signals from the front surface 110 of thescintillator array. Further, as shown in FIG. 22, when a ray source isnot used, and on a background of a natural radioactive ray 116, a signalcount N_(1b) and a signal count N_(2b) are calculated through adetermination calculation, in which the signal count N_(1b) isdetermined as the count of signals from the front surface 110 of thescintillator array, and the signal count N_(2b) is determined as thecount of signals from the rear surface 111 of the scintillator array.Then, (N₁−N_(1b))/(N₂−N_(2b)) and (N₂′−N_(2b))/(N₁′−N_(1b)) are defined.When (N₁−N_(1b))/(N₂−N_(2b)) and (N₂′−N_(2b))/(N₁′−N_(1b)) are equal andare the maximum, the parameters are defined as the optimal parameters.Further, a lead collimator 13 and a Ri ray source 114 are required toensure that the γ-ray is irradiated on only any one of the scintillatorarrays. The parameters decided in the above manner are pre-input to aprocessing circuit for the device during the stage in which the γ-raydetector single unit is installed in the actual PET device.

Patent Reference 1: Japanese Patent Publication No. H06-337289 (Page 2to 3 and FIG. 1)

Patent Reference 2: Japanese Patent Publication No. 2000-56023 (Page 2to 3 and FIG. 1)

However, the existing nuclear medical diagnostic device has thefollowing problem. In the case of the two-stage scintillator detector112, for example, having the scintillator arrays of the two-stagestructure as shown in FIG. 20, the parameters T₁, T₂, and K required bythe scintillator array identification mechanism are decided by theprocessing circuit for inspection during the inspection stage of theγ-ray detector 12 single unit, and these parameters are applied to theprocessing circuit for the device during the stage in which the γ-raydetector single unit is installed in the actual PET device. No matterthe processing circuit for inspection and the processing circuit for thedevice are manufactured based on the same specification, temperaturecharacteristics of a gain amplifier or other elements may be slightlydifferent, so as to generate individual differences. Therefore, theoptimal values of the parameters are not necessarily consistent, suchthat it is impossible to separate upper and lower parts, resulting inadverse impact on the image quality.

In another aspect, if it is intended to decide the parameters during thestage in which the γ-ray detector single unit is installed in the actualPET device, a large lead calibration jig and Ri ray source matching withthe PET device are required, so the operation is quite complicated.

SUMMARY OF THE INVENTION

In order to solve the above problems, according to one aspect of thepresent invention, it is to provide a nuclear medical diagnostic devicewhich adopts the following structure. That is, the nuclear medicaldiagnostic device includes a plurality of γ-ray detectors, a processingcircuit for the device, a frame, a parameter deciding mechanism foridentification, and an identification mechanism. Each of the γ-raydetectors is composed of a scintillator block, a light receivingelement, and an A/D converter, in which the scintillator block is formedby optically combining a plurality of scintillator arrays withluminescence pulses having different attenuation times in a γ-rayincident depth direction, the light receiving element converts theluminescence pulse emitted from the scintillator block to an electricsignal, and the A/D converter converts the electric signal output fromthe light receiving element, i.e., an analog signal, to a digitalsignal. The processing circuit for the device is used for calculating asignal count ratio according to the digital signals from the γ-raydetectors. The frame is for installing the plurality of γ-ray detectors.The parameter determining mechanism for identification is used forcalculating a first signal count ratio according to the digital signalobtained by irradiating the γ-ray on each scintillator array andobtaining a second signal count ratio by irradiating the γ-ray on thescintillator block and performing a measurement when the plurality ofγ-ray detectors are not respectively installed in the frame but areconnected to a processing circuit for inspection that is different fromthe processing circuit for the device, and obtaining a third signalcount ratio by irradiating the γ-ray on the respective scintillatorblock of the γ-ray detectors and performing a measurement when theplurality of γ-ray detectors are respectively installed in the frame andare connected to the processing circuit for the device, and deciding theparameters for identification according to the first signal count ratio,the second signal count ratio, and the third signal count ratio. Theidentification mechanism is used for identifying which scintillatorarray of the plurality of γ-ray detectors is irradiated by the γ-ray ofa measurement object according to the parameters for identification.

Further, according to one aspect of the present invention, the secondsignal count ratio and the third signal count ratio are obtained byirradiating the γ-ray on a front surface or a rear surface of thescintillator block and performing the measurement.

Further, according to one aspect of the present invention, the secondsignal count ratio and the third signal count ratio are obtained byirradiating the γ-ray on a lateral side of the scintillator block andperforming the measurement.

Further, according to one aspect of the present invention, a ray sourcefor transmission is used to irradiate the γ-ray for obtaining the thirdsignal count ratio.

Effect of the Invention

Through the above means, the optimal parameters may be decided. By usingthe optimal parameters, the position of a γ-ray source of themeasurement object may be accurately assigned, so as to provide ahigh-quality tomogram.

Further, by dividing the measurement into a measurement of the γ-raydetector single units and a measurement of the γ-ray detectors installedin the frame, the optimal parameters may be easily determined.

BRIEF DESCRIPTION OF THE DRAWINGS

The accompanying drawings are included to provide a furtherunderstanding of the invention, and are incorporated in and constitute apart of this specification. The drawings illustrate embodiments of theinvention and, together with the description, serve to explain theprinciples of the invention.

FIG. 1 is an outside view of a γ-ray detector of the present invention.

FIG. 2 is an illustrative view of a method for identifying ascintillator array of the present invention.

FIG. 3 is an illustrative view of an example of a position operatingcircuit of the γ-ray detector of the present invention.

FIG. 4 is an illustrative view of a position coding map of the γ-raydetector of the present invention.

FIG. 5 is an energy spectrum of the γ-ray detector of the presentinvention.

FIG. 6 is an illustrative view of a method for deciding parametersaccording to a first embodiment.

FIG. 7 is an illustrative view of the method for deciding the parametersaccording to the first embodiment.

FIG. 8 is an illustrative view of the method for deciding the parametersaccording to the first embodiment.

FIG. 9 is an illustrative view of the method for identifying thescintillator array according to the first embodiment.

FIG. 10 is a schematic view of a frame with a plurality of γ-raydetectors installed therein.

FIG. 11 is an illustrative view of the method for deciding parametersaccording to a second embodiment.

FIG. 12 is an illustrative view of the method for deciding theparameters according to the second embodiment.

FIG. 13 is an illustrative view of the method for deciding theparameters according to the second embodiment.

FIG. 14 is an illustrative view of the method for deciding theparameters according to a third embodiment.

FIG. 15 is an illustrative view of the method for deciding theparameters according to the third embodiment.

FIG. 16 is an illustrative view of a γ-ray detecting principle of anexisting γ-ray detector.

FIG. 17 is an illustrative view of a γ-ray detecting principle of anexisting DOI γ-ray detector.

FIG. 18 is a waveform diagram of an electrical signal output from theexisting DOI γ-ray detector.

FIG. 19 shows values obtained by integrating the data of a time seriesoutput from the existing DOI γ-ray detector.

FIG. 20 is an illustrative view of the method for deciding theparameters of an existing nuclear medical diagnostic device.

FIG. 21 is an illustrative view of the method for deciding theparameters of the existing nuclear medical diagnostic device.

FIG. 22 is an illustrative view of the method for deciding theparameters of the existing nuclear medical diagnostic device.

DESCRIPTION OF THE EMBODIMENTS

Reference will now be made in detail to the present embodiments of theinvention, examples of which are illustrated in the accompanyingdrawings. Wherever possible, the same reference numbers are used in thedrawings and the description to refer to the same or like parts.

First Embodiment

The detailed structure of a γ-ray detector according to the firstembodiment of the present invention is shown in the drawings and isillustrated in the following. FIG. 1 is an outside view of a γ-raydetector 10 having scintillator arrays of a two-stage structure of thepresent invention. As shown in FIG. 1, in the γ-ray detector 10, ascintillator block 1 is arranged by being divided in a γ-ray incidentdepth direction, that is, the radiation detector 10 is a depth ofinteraction (DOI) γ-ray detector having three-dimensionally arrangedscintillators. The DOI γ-ray detector of this embodiment hasscintillator arrays of a two-stage structure.

The γ-ray detector 10 of this embodiment is approximately formed by fourparts. The first part is a scintillator array 11F having scintillators1SF that are two-dimensionally and compactly arranged and have aluminescence pulse with a shorter attenuation time. The scintillators1SF are divided by appropriately sandwiching a light reflective material12, and 64 scintillators 1SF are arranged in a manner of having eightscintillators in the X direction and eight scintillators in the Ydirection. The second part is a scintillator array 11R havingscintillators 1SR that are two-dimensionally and compactly arranged andhave a longer attenuation time of the luminescence pulse. Thescintillators 1SR are divided by appropriately sandwiching the lightreflective material 12 there-between, and 64 scintillators 1SR arearranged in a manner having eight scintillators in the X direction andeight scintillators in the Y direction. The scintillator array 11F andthe scintillator array 11R form the scintillator block 1. The third partis a light guide 20, which is optically combined with the scintillatorblock 1, includes embedded lattice frames combined with a lightreflective material 13 (not shown), and is divided into a plurality ofsmall areas. The fourth part is four photomultipliers 31, 32, 33, and 34respectively to optically combine with the light guide 20.

The scintillators 1SF having the luminescence pulse with a shorterattenuation time employ, for example, inorganic crystals, such asGd₂SiO₅:Ce1.5 mol % (GSO:Ce1.5 mol %), Gd₂SiO₅:Ce1.5 mol % (GSOZ:Ce1.5mol %) doped with Zr, Lu₂SiO₅:Ce(LSO), LuYSiO₅:Ce(LYSO), LaBr₃:Ce,LaCl₃:Ce, and LuI:Ce. In another aspect, the scintillators 1SR havingthe luminescence pulse with a longer attenuation time employ inorganiccrystals, such as Gd₂SiO₅:Ce0.5 mol % (GSO:Ce0.5 mol %), Gd₂ SiO₅:Ce0.5mol % (GSOZ:Ce0.5 mol %) doped with Zr, Bi₄Ge₃O₁₂ (BGO), andLu_(0.4)Gd_(1.6)SiO₅:Ce (LGSO).

Table 1 lists data of the attenuation time of each scintillator.

TABLE 1 Attenuation Time Scintillator [ns] Gd₂SiO₅:Ce1.5 mol %(GSO:Ce1.5 mol %) 40 Gd₂SiO₅:Ce1.5 mol % (GSOZ:Ce1.5 mol %) 40 dopedwith Zr Lu₂SiO₅:Ce (LSO) 40 LuYSiO₅:Ce (LYSO) 40 LaBr₃:Ce 27 LaCl₃:Ce 70LuI:Ce 25 Gd₂SiO₅:Ce0.5 mol % (GSO:Ce0.5 mol %) 80 Gd₂ SiO₅:Ce0.5 mol %(GSOZ:Ce0.5 mol %) 80 doped with Zr Bi₄Ge₃O₁₂ (BGO) 300Lu_(0.4)Gd_(1.6)SiO₅:Ce (LGSO) 43

The scintillator block 1 is formed by optically combining the twoscintillator arrays 11F and 11R having luminescence pulses withdifferent attenuation times in the γ-ray incident depth direction (Zdirection). The scintillator array 11F has a plurality of scintillators1SF that is two-dimensionally and compactly arranged and has aluminescence pulse with a shorter attenuation time, and the scintillatorarray 11R has a plurality of scintillators 1SR that is two-dimensionallyand compactly arranged and has a luminescence pulse with a longerattenuation time. Particularly, in the scintillator block 1, forexample, Gd₂SiO₅:Ce1.5 mol % (GSO:Ce1.5 mol %) is used as thescintillators 1SF having the luminescence pulse on a γ-ray incident side(front segment) with a shorter attenuation time, and Gd₂SiO₅:Ce0.5 mol %(GSO:Ce0.5 mol %) is used as the scintillators 1SR having theluminescence pulse on the light guide 20 side (back segment) with alonger attenuation time.

The two scintillator arrays 11F and 11R are respectively formed by 8×8(in the X direction and Y direction) chip-shaped scintillators, and thelight reflective material 12, a light transmissive material (not shown),or an optical binding agent (not shown), which enables the lightgenerated when the γ-ray is incident to distribute in the X directionand Y direction according to a ratio, are sandwiched or filled in somepositions between the chip-shaped scintillators.

The light guide 20 guides the light generated by the scintillators 11Fand 11R of the scintillator block 1 to the photomultipliers 31˜34. Thelight guide 20 is sandwiched between the scintillator block 1 and thephotomultipliers 31˜34 in an optically combining manner by using theoptical binding agent.

The light generated by the scintillator arrays 11F and 11R is incidenton photomultiplier photoelectric conversion films on four sides, and iselectronically amplified, and then is finally converted to electricsignals (analog signals) for output. Therefore, the outputs of thephotomultipliers 31˜34 become the output of the γ-ray detector 10.

Here, the light in the scintillator block 1 is guided to thephotomultipliers 31˜34 by the optically combined light guide 20. At thistime, the position, length, and angle of each light reflective material13 (not shown) in the light guide 20 are adjusted, such that the outputratio of the photomultiplier 31 (33) and the photomultiplier 32 (34)arranged in the X direction is changed according to a fixed proportion.

Here, in the present invention, the parameters T1, T2, and K required bythe scintillator array identification mechanism are decided in thefollowing manner. As shown in FIG. 2, a Ri ray source 35 irradiates fromthe front the γ-ray on the γ-ray detector 10 having scintillator arraysof the two-stage structure and disposed in the dark box 15, and aposition coding map and an energy spectrum are measured. That is, if theoutput of the photomultiplier 31 is set to P1, the output of thephotomultiplier 32 is set to P2, the output of the photomultiplier 33 isset to P3, and the output of the photomultiplier 34 is set to P4, acalculated value {(P1+P3)−(P2+P4)}/(P1+P2+P3+P4) representing a positionin the X direction is obtained, and a calculated value{(P1+P2)−(P3+P4)}/(P1+P2+P3+P4) representing a position in the Ydirection is also obtained.

FIG. 3 is a block diagram of a structure of a position operating circuitof the γ-ray detector 10. The position operating circuit is formed byadders 71, 72, 73, and 74, and position recognizing circuits 75 and 76.As shown in FIG. 3, in order to detect the incident position of theγ-ray in the X direction, the output P1 of the photomultiplier 31 andthe output P3 of the photomultiplier 33 are input to the adder 71, andthe output P2 of the photomultiplier 32 and the output P4 of thephotomultiplier 34 are input to the adder 72. Each added output (P1+P3)and (P2+P4) of the two adders 71 and 72 are input to the positionrecognizing circuit 75. According to the two added outputs, the incidentposition of the γ-ray in the X direction is obtained. Similarly, inorder to detect the incident position of the γ-ray in the Y direction,each added output (P1+P2) and (P3+P4) are input to the positionrecognizing circuit 76. According to the two added outputs, the incidentposition of the γ-ray in the Y direction is obtained. According to thepositions of the γ-ray incident on the scintillators, the calculatedresults obtained above are represented by the position coding mapshowing the position recognition information, as shown in FIG. 4.

In another aspect, the calculated value (P1+P2+P3+P4) represents theenergy relative to the event, and is calculated as an energy spectrum.As an example, FIG. 5 is an energy spectrum relative to a representingportion 80 on the position coding map. Here, the luminescence outputs ofthe two scintillators are different, so two energy peak values PF and PRoccur. In this embodiment, PF is equivalent to Gd₂SiO₅: Ce1.5 mol %(GSO: Ce1.5 mol %) and PR is equivalent to Gd₂SiO₅: Ce0.5 mol % (GSO:Ce0.5 mol %).

For the γ-ray detector 10 having the scintillator arrays of thetwo-stage structure, the method for deciding the parameters T₁, T₂, andK required by the scintillator array identification mechanism isillustrated. As shown in FIG. 6, during a first inspection stage of theγ-ray detector single unit, the γ-ray detector 10 is disposed in thedark box 15, and initial values of the parameters are input to aprocessing circuit for inspection 16. In this state, the γ-ray from theRi ray source 35 calibrated by the lead collimator 36 only irradiatesonto the scintillator array 11F. According to the input parameters, inthe processing circuit for inspection 16, a signal count N₁ and a signalcount N₂ are calculated through a determination calculation, in whichthe signal count N1 is judged as the count of signals from thescintillator array 11F, and the signal count N2 is judged as the countof signals from the scintillator array 11R.

Then, as shown in FIG. 7, the γ-ray from the Ri ray source 35 calibratedby the lead collimator 36 only irradiates onto the scintillator array11R. According to the input parameters, in the processing circuit forinspection 16, a signal count N₂′ and a signal count N₁′ are calculatedthrough a determination calculation, in which the signal count N2′ isdetermined as the count of signals from the scintillator array 11R, andthe signal count N1′ is determined as the count of signals from thescintillator array 11F.

Further, as shown in FIG. 8, when the ray source is not used, mainly thecount caused by a natural radioactive ray 17 is measured. On the stateof a background, in the processing circuit for inspection 16, a signalcount N_(1b) and a signal count N_(2b) are calculated through adetermination calculation, in which the signal count N_(1b) isdetermined as the count of signals from the scintillator array 11F, andthe signal count N_(2b) is determined as the count of signals from thescintillator array 11R.

Here, R₁=(N₁−N_(1b))/(N₂−N_(2b)) and R₁′=(N₂′−N_(2b))/(N₁′−N_(1b)) aredefined as a first signal count ratio. When the values of R and R′ areequal and become the maximum, the parameters T₁, T₂, and K are definedas the optimal parameters.

Then, as shown in FIG. 9, during a second inspection stage of the γ-raydetector single unit, the γ-ray detector 10 is disposed in the dark box15, the optimal parameters T₁, T₂, and K decided above are input to theprocessing circuit for inspection 16. In this state, the Ri ray source35 is arranged at a distance d relative to the γ-ray detectors 10, andirradiates the γ-ray on the front surface of the γ-ray detectors 10. Thedistance d is equal to a distance from the center of the PET device tothe surface of the γ-ray detector. According to the input parameters, inthe processing circuit for inspection 16, a signal count N_(F) and asignal count N_(R) are calculated through the determination calculation,in which the signal count N_(F) is determined as the count of signalsfrom the scintillator array 11F, and the signal count N_(R) isdetermined as the count of signals from the scintillator array 11R.Here, R₂=N_(F)/N_(F) is defined as a second signal count ratio. Untilnow, it is the inspection stage of the γ-ray detector single unit.

As shown in FIG. 10, during a third inspection stage, the plurality ofγ-ray detectors 10 is installed in a frame (not shown) according to theactual number of the formed PET devices. Each γ-ray detector 10 isconnected to a processing circuit for the device 18. FIG. 10 only showsthe γ-ray detectors disposed along ¼ of the entire circumference, but inreality, the γ-ray detectors are present along the entire circumference,and the γ-ray detectors are accommodated in an appropriate dark box (notshown). In addition, it is preferred that the frame actually forming thePET device is being used as the frame; but, a dedicated frame forinspection may also be used.

No matter the processing circuit for inspection 16 and the processingcircuit for the device 18 used during the inspection stage aremanufactured according to the same specification, temperaturecharacteristics of a gain amplifier or other elements may be slightlydifferent, so as to generate individual differences. Therefore, theoptimal parameters T₁, T₂, and K determined above are not necessarilyconsistent, such that the optimal parameters must be decided again, butthe optimal parameters T₁, T₂, and K determined above are temporarilyused as the initial values and are input to the processing circuit forthe device 18. Then, as shown in FIG. 10, a Ri ray source 37 is arrangedon the center position of the PET device at the distance d relative toall the γ-ray detectors 10, and irradiates the γ-ray on the frontsurfaces of all the γ-ray detectors 10.

According to the input parameters, in the processing circuit for thedevice 18, a signal count N_(F)′ and a signal count N_(R)′ arecalculated through the determination calculation, in which the signalcount N_(F)′ is determined as the count of signals from the scintillatorarray 11F, and the signal count N_(R)′ is judged as the count of signalsfrom the scintillator array 11R.

In this invention, R₃=N_(F)′/N_(R)′ is defined as a third signal countratio. Therefore, the condition of the parameters that enables thesecond signal count ratio R₂=N_(F)/N_(R) obtained above and the thirdsignal count ratio R₃=N_(F)′/N_(R)′ calculated this time being equal isidentified, and the parameters are defined as optimal parameter valuesT′₁, T₂′, and K′ required by the scintillator array identificationmechanism.

As described above, the optimal parameter values T′₁, T₂′, and K′ aredecided by making the second signal count ratio in the inspection stageand the third signal count ratio of the PET device be equal, such thatthe optimal parameter values may be determined correctly.

Second Embodiment

The detailed structure of the γ-ray detector according to the secondembodiment of the present invention is shown in the drawings and isillustrated in the following. This embodiment is used for the situationin which the optimal values are more strictly decided. The scintillatorarray identification mechanism defines the first signal count ratio,which is R₁=(N₁−N_(1b))/(N₂−N_(2b)) and R₁′=(N₂′−N_(2b))/(N₁′−N_(1b)),obtained during the first inspection stage of the γ-ray detector singleunit, and defines the parameters T₁, T₂, and K when the values of R andR′ are equal and become the maximum as the optimal values. Until now,the process is the same as that of the first embodiment (FIGS. 6 to 8).

As shown in FIG. 11, during a second inspection stage of the γ-raydetector single unit, the γ-ray detector 10 is disposed in the dark box15, the optimal parameters T₁, T₂, and K decided above are input to theprocessing circuit for inspection 16. In this state, the Ri ray source35 is arranged on a position at a distance d relative to the γ-raydetectors 10, and irradiates the γ-ray on the front surfaces of theγ-ray detectors 10. The distance d is equal to a distance from thecenter of the PET device to the surface of the γ-ray detectors.According to the input parameters, in the processing circuit forinspection 16, a signal count N_(F) and a signal count N_(R) arecalculated through the judging calculation, in which the signal countN_(F) is judged as the count of signals from the scintillator array 11F,and the signal count N_(R) is judged as the count of signals from thescintillator array 11R. In this invention, R₂=N_(F)/N_(F) is defined asa second signal count ratio.

Further, i as shown in FIG. 12, the Ri ray source 35 is arranged on aposition at a distance d′ relative to the γ-ray detectors 10, andirradiates the γ-ray on the rear surfaces of the γ-ray detectors 10. Thedistance d′ is equal to a distance from the rear surfaces of the γ-raydetectors 10 to the Ri ray source when the γ-ray detectors are installedin the PET device. According to the input parameters, in the processingcircuit for inspection 16, a signal count N_(Fb) and a signal countN_(Rb) are calculated through the determination calculation, in whichthe signal count N_(Fb) is determined as the count of signals from thescintillator array 11F, and the signal count N_(Rb) is determined as thecount of signals from the scintillator array 11R. In this invention,R_(2b)=N_(Fb)/N_(Rb) is defined as the second signal count ratio. Untilnow, it is the inspection stage of the γ-ray detector single unit.

Then, as shown in FIG. 13, during a third inspection stage, theplurality of γ-ray detectors 10 is installed in a frame (not shown)according to the actual number of the formed PET devices. Each γ-raydetector 10 is connected to a processing circuit for the device 18. FIG.13 only shows the γ-ray detectors disposed along a ¼ of the entirecircumference, but in reality, the γ-ray detectors are disposed alongthe entire circumference, and the γ-ray detectors are accommodated in anappropriate dark box (not shown). In addition, it is preferred that theframe actually forming the PET device is being used as the frame; but, adedicated frame for inspection may also be used.

No matter the processing circuit for inspection 16 and the processingcircuit for the device 18 used during the inspection stage aremanufactured according to the same specification, temperaturecharacteristics of a gain amplifier and other elements may be slightlydifferent, so as to generate individual differences. Therefore, theoptimal parameters T₁, T₂, and K decided above are not necessarilyconsistent, such that the optimal parameters must be decided again, butthe optimal parameters T₁, T₂, and K decided above are temporarily usedas the initial values and are input to the processing circuit for thedevice 18. Then, as shown in FIG. 13, an Ri ray source 37 is arranged onthe center position of the PET device at the distance d relative to allthe γ-ray detectors 10, and irradiates the γ-ray on the front surfacesof all the γ-ray detectors 10.

According to the input parameters, in the processing circuit for thedevice 18, a signal count N_(F)′ and a signal count N_(R)′ arecalculated through the determination calculation, in which the signalcount N_(F)′ is determined as the count of signals from the scintillatorarray 11F, and the signal count N_(R)′ is judged as the count of signalsfrom the scintillator array 11R. Here, R₃=N_(F)′/N_(F)′ is defined as athird signal count ratio. Therefore, the condition of the parametersthat enables the second signal count ratio R₂=N_(F)/N_(R) obtained aboveand the third signal count ratio R₃=N_(F)′/N_(R)′ calculated this timebeing equal is identified, and the parameters are defined as optimalparameter values T′₁, T₂′, and K′ required by the scintillator arrayidentification mechanism.

Further, as shown in FIG. 13, for all the γ-ray detectors 10, an Ri raysource 38 (although the Ri ray source 37 is shown, here the Ri raysource 37 does not exist) is arranged on a position of the PET device atthe distance d′ relative to all the γ-ray detectors 10, and irradiatessequentially the γ-ray on the rear surfaces of all the γ-ray detectors10.

According to the input parameters, in the processing circuit for thedevice 18, a signal count N_(Fb)′ and a signal count N_(Rb)′ arecalculated through the determination calculation, in which the signalcount N_(Fb)′ is determined as the count of signals from thescintillator array 11F, and the signal count N_(Rb)′ is determined asthe count of signals from the scintillator array 11R. In this invention,R₃=N_(Fb)′/N_(Rb)′ is defined as the third signal count ratio.Therefore, the condition of the parameters that enables the secondsignal count ratio R₂=N_(Fb)/N_(Rb) obtained above and the third signalcount ratio R₃=N_(Fb)′/N_(Rb)′ calculated this time being equal isidentified, and the parameters are defined as optimal parameter valuesT″₁, T₂″, and K″ required by the scintillator array identificationmechanism.

Moreover, medium values of the optimal parameter values T′₁, T₂′, and K′and the optimal parameter values T″₁, T₂″, and K″ are respectively used,such that the optimal parameter values may be decided more correctly.

Third Embodiment

The detailed structure of the γ-ray detector according to the thirdembodiment of the present invention is shown in the drawings and isillustrated in the following. The scintillator array identificationmechanism defines the first signal count ratio, which isR₁=(N₁−N_(1b))/(N₂−N_(2b)) and R₁′=(N₂′−N_(2b))/(N₁′−N_(1b)), obtainedduring the first inspection stage of the γ-ray detector single unit, anddefines the parameters T₁, T₂, and K when the values of R and R′ areequal and become the maximum as the optimal values. Until now, theprocess is the same as that of the first embodiment (FIGS. 6 to 8).

Then, as shown in FIG. 14, during a second inspection stage of the γ-raydetector single unit, the γ-ray detector 10 is disposed in the dark box15, the optimal parameters T₁, T₂, and K decided above are input to theprocessing circuit for inspection 16. In this state, the Ri ray source35 is arranged on a position at a distance 1 relative to the γ-raydetectors 10, and irradiates the γ-ray on a lateral side of the γ-raydetectors 10. The distance 1 is equal to a distance from a ray sourcefor transmission on the PET device to the surface of the γ-ray detector.According to the input parameters, in the processing circuit forinspection 16, a signal count N_(Fc) and a signal count N_(Rc) arecalculated through the determination calculation, in which the signalcount N_(Fc) is determined as the count of signals from the scintillatorarray 11F, and the signal count N_(Rc) is determined as the count ofsignals from the scintillator array 11R. In this invention,R₂=N_(Fc)/N_(Rc) is defined as a second signal count ratio. Until now,it is the inspection stage of the γ-ray detector single unit.

Then, as shown in FIG. 15, during a third inspection stage, theplurality of γ-ray detectors 10 is installed in a frame (not shown)according to the actual number of the formed PET devices. Each γ-raydetector 10 is connected to a processing circuit for the device 18. FIG.15 only shows the γ-ray detectors disposed along ¼ of the entirecircumference; but in reality, the γ-ray detectors are present along theentire circumference, and the γ-ray detectors are accommodated in anappropriate dark box (not shown). In addition, it is preferred that theframe actually forming the PET device is being used as the frame; but adedicated frame for inspection may also be used. Further, no matter theprocessing circuit for inspection 16 and the processing circuit for thedevice 18 used during the inspection stage are manufactured according tothe same specification, temperature characteristics of a gain amplifierand other elements may be slightly different so as to generateindividual differences. Therefore, the optimal parameters T₁, T₂, and Kdecided above are not necessarily consistent, such that the optimalparameters must be decided again. However, the optimal parameters T₁,T₂, and K decided above are temporarily used as the initial values andinput to the processing circuit for the device 18.

As shown in FIG. 15, a rotating ray source for transmission 39 isarranged on a position at the distance 1 relative to all the γ-raydetectors 10. In addition, in order to obtain absorption correctiondata, the ray source for transmission 39 and detectors for transmission40 are usually arranged in the PET device. Accordingly, the ray sourcefor transmission 39 may irradiate sequentially the γ-ray on thedetectors for transmission 40 through a rotating mechanism 42. The raysource for transmission 39 is arranged in a lead case 41, and has anopenable window used to irradiate the γ-ray detectors 10 and an openablewindow used to irradiate the detector for transmission 40. The raysource for transmission 39 may irradiate sequentially the γ-ray on thelateral sides of the γ-ray detectors 10 through the rotating mechanism42.

According to the input parameters, in the processing circuit for thedevice 18, a signal count N_(Fc)′ and a signal count N_(Rc)′ arecalculated through the determination calculation, in which the signalcount N_(Fc)′ is determined as the count of signals from thescintillator array 11F, and the signal count N_(Rc)′ is determined asthe count of signals from the scintillator array 11R. In this invention,R₃=N_(Fc)′/N_(Rc)′ is defined as a third signal count ratio. Therefore,the condition of the parameter that enables the second signal countratio R₂=N_(Fc)/N_(Rc) obtained above and the third signal count ratioR₃=N_(Fc)′/N_(Rc)′ calculated this time being equal is identified, andthe parameters are defined as optimal parameter values T′₁, T₂′, and K′required by the scintillator array identification mechanism.

In the nuclear medical diagnostic device of the present invention, byadopting the above method, the optimal parameter values may be correctlyand simply decided without using special external ray sources. A highimage quality is thereby maintained, such that the high image quality isachieved through a simple operation. Further, after the γ-ray detectorsare installed in the actual PET device, the large lead calibration jigmatching with the PET device is not required, so the operation is quiteeasy.

The present invention is not limited to the embodiments above, and mayalso be implemented in the following variations.

In the embodiments, the PET device is taken as an example forillustration. However, the present invention may be applied to othernuclear medical diagnostic devices besides the PET device, as long asthe devices may at the same time count the radioactive rays emitted fromthe detected body on which the radioactive agent is applied.

The embodiments of the present invention may be applied in a deviceformed by combining the nuclear medical diagnostic device and an X-rayCT device as a PET-CT having the PET device and the X-ray CT device.

In the embodiments, the scintillator block 1 is farmed by combining thetwo (layers of) scintillator arrays 11F and 11R, but may also be formedby combining a plurality of (layers of) scintillator arrays besides two(layers of) scintillator arrays. Further, the number of thescintillators in each of the scintillator arrays 11F and 11R is 8×8, butmay also be multiple besides 8×8.

In the embodiments, the light receiving elements are thephotomultipliers 31˜34, but other light receiving elements may also beused, for example, photodiodes or avalanche photodiodes.

INDUSTRIAL AVAILABILITY

As described above, the present invention is suitable for medical andindustrial radiation imaging devices.

It will be apparent to those skilled in the art that variousmodifications and variations can be made to the structure of the presentinvention without departing from the scope or spirit of the invention.In view of the foregoing, it is intended that the present inventioncover modifications and variations of this invention provided they fallwithin the scope of the following claims and their equivalents.

1. A nuclear medical diagnostic device, comprising: a plurality of γ-raydetectors, wherein each of the γ-ray detectors is composed of ascintillator block, a light receiving element, and an A/D converter, thescintillator block is formed by optically combining a plurality ofscintillator arrays with luminescence pulses in a γ-ray incident depthdirection having different attenuation times, the light receivingelement converts the luminescence pulse emitted from the scintillatorblock to an electric signal, and the A/D converter converts the electricsignals output from the light receiving element, which are analogsignals, to digital signals; a processing circuit for the device, forcalculating a signal count ratio according to the digital signals fromthe γ-ray detectors; a frame, for installing the plurality of γ-raydetectors; a parameter deciding mechanism for identification,calculating a first signal count ratio according to the digital signalsobtained by irradiating the γ-ray on each of the plurality of thescintillator arrays and obtaining a second signal count ratio byirradiating the γ-ray on the scintillator block and performing ameasurement when the plurality of γ-ray detectors are not respectivelyinstalled in the frame but are connected to a processing circuit forinspection that is different from the processing circuit for the device,and obtaining a third signal count ratio by irradiating the γ-ray on therespective scintillator block of the γ-ray detectors and performing ameasurement when the plurality of γ-ray detectors are respectivelyinstalled in the frame and are connected to the processing circuit forthe device, and determining parameters for identification according tothe first signal count ratio, the second signal count ratio, and thethird signal count ratio; and an identification mechanism, foridentifying which one of the plurality of the scintillator arrays of theplurality of γ-ray detectors is irradiated by the γ-ray of a measurementobject according to the parameters for identification.
 2. The nuclearmedical diagnostic device according to claim 1, wherein the secondsignal count ratio and the third signal count ratio are obtained byirradiating the γ-ray on a front surface or a rear surface of thescintillator block and performing the measurement.
 3. The nuclearmedical diagnostic device according to claim 1, wherein the secondsignal count ratio and the third signal count ratio are obtained byirradiating the γ-ray on a lateral side of the scintillator block andperforming the measurement.
 4. The nuclear medical diagnostic deviceaccording to claim 3, wherein a ray source for transmission is used toirradiate the γ-ray for obtaining the third signal count ratio.